In situ Forming Injectable Hydrogels for Drug Delivery and Wound Repair

Abstract

Hydrogels have been utilized in regenerative applications for many decades because of their biocompatibility and similarity in structure to the native extracellular matrix. Initially, these materials were formed outside of the patient and implanted using invasive surgical techniques. However, advances in synthetic chemistry and materials science have now provided researchers with a library of techniques whereby hydrogel formation can occur in situ upon delivery through standard needles. This provides an avenue to minimally invasively deliver therapeutic payloads, fill complex tissue defects, and induce the regeneration of damaged portions of the body. In this review, we highlight these injectable therapeutic hydrogel biomaterials in the context of drug delivery and tissue regeneration for skin wound repair.

Graphical abstract

1. Introduction

1.1. Injectable Gels as Therapeutic Agents

Hydrogels are highly-crosslinked, water swollen networks of hydrophilic polymers, which have been studied extensively over the past six decades, and have demonstrated profound promise as bio-compatible materials in numerous therapeutic applications¹. These materials can be derived from both natural and synthetic sources². Naturally occurring polymers such as chitosan, alginate, hyaluronic acid (HA), collagen, and gelatin are inherently biodegradable and often come pre-functionalized with integrin binding sites allowing for adhesion and coordinated cellular responses. Unfortunately, the utilization of these materials is limited due to significant batch-to-batch variability and potential immunogenicity within foreign hosts. In contrast, synthetic polymers such as poly(ethylene glycol) (PEG), polyacrylamide (PAM), poly(vinyl alcohol) (PVA), and poly(methyl methacrylate) (PMMA) are appealing due to their strong mechanical properties, tailorable structure and low immunogenicity, but lack innate bio-functionality and must undergo significant post-processing in order to elicit desired responses in vivo. More complex, hydrogel systems have also been developed to circumvent the limitations presented through designing scaffolds from a single polymer backbone. These materials come in the form of either co-polymers³, where multiple backbone groups are crosslinked together, or inter-penetrating networks (IPNs)⁴, where a polymer mesh is constructed from the binding of oligomer chains within an already assembled polymeric scaffold. In this manner, hydrogel materials may be precisely modified to highlight the optimal properties of each of their constituent components, resulting in an even greater degree of control towards regenerative outcomes.

Historically, hydrogels were pre-formed and delivery of these materials to target sites in patients necessitated the use of highly invasive surgical procedures. However, influential work in the late 90's demonstrated that hydrogel precursors could be injected through a standard syringe and crosslinked locally through transdermal light-induced photopolymerization⁵. Nowadays, minimally invasive delivery of hydrogels through injection has gained significant traction in the biomedical community. Injectable materials have several inherent advantages over their pre-formed counterparts. In short, associated implantation procedures are lower cost, with patient discomfort significantly reduced after delivery, delicate therapeutic materials dissolved within the materials are shielded from injection associated shear forces⁶ and can be released with complex dynamics⁷, and lastly tissue regeneration is aided by the ability of these materials to mold into the shape of the injection cavity⁸, allowing for universal off the shelf treatment within any non-standard geometry. Continued advances in our understanding of polymer chemistry have fostered the development of numerous biomaterials which can be injected as viscous liquids and subsequently solidified through variations in their local microenvironment (temperature^9,10, pH¹⁰, ion concentration¹⁰), application of an external stimulus (light¹¹), or affinity based self-organization in the case of peptides^12,13 and other physically associating functional moieties¹⁴ (Table 1). This diversity in injectable hydrogel technologies is critical for the recapitulation of complex extracellular environments, organization of cellular behavior, and adequate delivery of therapeutic small molecules. Successfully blending components from these systems will enable the development and optimization of novel therapeutic injectable hydrogels.

Table 1. Examples of materials used for drug delivery and tissue regeneration.

Hydrogel Component	Application	Refs.
Local Microenvironment
Temperature Driven
poly(D,L-lactide-co-glycolide) (PLGA)/PEG triblock copolymers (PLGA-PEG-PLGA)	Drug Delivery	51
polyethylene glycol-poly(L-alanine) (PEG-PLA)	Cell Scaffold	56
poly(N-isopropyl acrylamide)	Cell Scaffold	58
soluble ECM/methylcellulose	Cell Scaffold	59
pH Driven
PEG-diacrylate (PEGDA), acrylic acid and alginate	Wound Dressing	68
Ionic Concentration Driven
alginate/multi walled carbon nanotubes	Cell Scaffold	71
alginate/PEG/hyaluronic acid	Cell Scaffold/Drug Delivery	75
alginate/PEG	Cell Scaffold	79
Self Assembly
Peptide
RADA16	Cell Scaffold/Drug Delivery	86, 88-90
Fmoc dipeptides	Cell Scaffold	91, 92
Nap-GFFYGGGWRESAI/TIP-1 crosslinker	Cell Scaffold/Drug Delivery	93
Leucine-α/β-dehydrophenylalanine	Drug Delivery	94
Covalently Bonded
Photo-Initiated
gelatin-methacrylate	Cell Scaffold	108, 109
gelatin-methacrylate/HA-methacrylate	Cell Scaffold	110
Reactive Precursors
8-arm PEG cysteine/N-hydroxysuccinimide	Drug Delivery	122
carboxymethyl chitosan/dextran	Cell Scaffold	124
konjac glucomannan-tyramine/heparin-tyramine	Cell Scaffold/Cytokine Sequestration	126

In this review, we seek to highlight advances in the design and development of injectable hydrogel materials towards application in skin. We begin by giving a brief, high level overview of skin biology and explain how deviations in responses during the wound healing cascade can lead to the formation chronic wounds. We next describe and motivate two frequent uses of injectable hydrogels as scaffolds for tissue regrowth and as depots for the release of small molecule or cellular therapeutics. Lastly, we devote the bulk of this review to a description of various injectable hydrogel systems and their wound healing applications. We start with the most basic systems derived from materials whose gelation is induced by simple environmental changes within the site of injection, and build towards newer systems such as guest-host mediated shear-thinning hydrogels, cryogels, and microporous annealed particle gels.

1.2. Rescuing Aberrant Skin Properties with Injectable Hydrogels

The skin is a highly organized, multi-faceted organ which serves as the primary line of defense for the human body. At the most basic level of classification, skin can be broken down into three main layers, each with unique properties that prove critical to its physiology (Fig 1a). The outermost layer, the epidermis, is roughly 50 to 100 cell layers thick and is mainly composed of melanocytes and senescent keratinocytes which provide protection against pathogens, UV radiation and mechanical stresses through their production of melanin and keratin respectively¹⁵. The dermis sits below the epidermis and is comprised of a complex network of structural proteins and proteoglycans which impart mechanical integrity to the overall tissue. Additionally, the dermis plays host to many higher order structures (sebaceous and sweat glands, hair follicles, and arrector pili muscles) and processes (oxygen exchange, nerve signaling) which prove imperative in maintaining cellular nourishment, regulating temperature homeostasis and responding to external stimuli^16,17. Lastly, the lowermost, subcutaneous layer of skin is mainly utilized as a depot for stored fat but also plays important roles by linking the more superficial layers to underlying muscle and bone¹⁶.

Figure 1. Skin biology and wound recovery.

(A) Skin is composed of three primary layers. The outermost layer, the epidermis, functions as a protective barrier and limits the interaction of underlying tissues with damaging pathogens, high energy UV radiation, and external forces. Below the epidermis, the dermis imparts mechanical integrity onto the tissue while facilitating higher order functions including oxygen exchange, temperature regulation, sweat production, nerve signaling, and hair growth. Finally, subcutaneous tissue serves as a depot for stored fats and anchors the more superficial layers of skin to underlying bone and muscle. (B) Upon injury, the skin progresses through a complex yet finely regulated wound healing response. Immediately following damage, platelets from the blood stream accumulate at the wound surface and form a fibrin clot which halts blood loss and simultaneously protects the injured area from foreign species. Macrophages and neutrophils are next recruited to the area and degrade pathogens to prime the area for recovery, while hypoxic signaling factors induce the formation of new vasculature. The wound is then healed through successive iterations of cellular migration, ECM maturation, and myofibroblast driven epidermal closure. (C) When wounds become chronic the normal healing response is impaired and proper regeneration can take months or even years to occur. Chronicity is driven by an imbalance between the rates of immune cell mediated degradation and the regeneration of functional ECM. Injectable hydrogel biomaterials offer an attractive option for treating chronic wounds because of their inherent ability to fill the wound defect while mimicking natural ECM, both by providing a scaffold for tissue ingrowth and by controlling the availability of regenerative signaling molecules.

Unsurprisingly, damage to the skin is fairly common. In most scenarios, regeneration is not difficult and takes place through a linear progression of overlapping events comprising initial immune response, tissue remodeling, cellular proliferation and maturation¹⁸ (Fig 1b). Unfortunately, there are many instances where such coordinated response is not feasible. When damage is extensive and penetrates deep into the dermis or sub-dermal layers as observed in 2^nd and 3^rd degree burns much of the bulk tissue organization is lost and regenerative cues are either completely absent or deregulated resulting in the production of highly fibrotic scar tissue¹⁹. Similarly, metabolic deficiencies arising from common diseases such as diabetes or venous insufficiency can lead to hyperinflammatory states which degrade the extracellular matrix (ECM) and diminish the availability of signaling factors and their associated cell surface receptors in the wound bed^20,21. These effects induce the transformation of many acute wounds into painful chronic ulcers which may take months or years to heal causing significantly reduced quality of life and potentially resulting in amputation if ignored or improperly treated²².

Hydrogels offer the unique capability to serve as tissue templates, where critical components of non-damaged tissue are temporarily recovered within the site of injury until regeneration and re-regulation of the wounded area concludes (Fig 1c). Ideally, these materials would serve as single application treatments, whereby they sequester damaging reactive oxygen species (ROS), resupply necessary growth signals, provide an environment conducive to cellular attachment and proliferation and degrade at a time frame longer than that necessary for proper healing. Unfortunately, successful engineering of materials that exhibit all of these properties remains difficult. There are currently many skin substitutes available on the market for clinical use. However, many of these materials are formed from reconstituted ECMs and must undergo numerous processing steps in order to remain immunogenically inert when transferring from xenogeneic or allogeneic sources. Furthermore, once formed, these artificial skins are physically homogeneous simple structures which serve as protective barriers but do not allow for any higher order skin processes to take place and are usually significantly undernourished due to decreased vascular permeability²³. Newly developed synthetic and semi-synthetic hydrogels offer more control over the spatial variation in their physical properties and encapsulated cargo concentrations²⁴ offering hope that more advanced products will soon reach the clinic, but even these materials still leave room for growth. Thus, the optimization of hydrogel design will rely heavily upon advances in basic science and our ability to replicate the complex dynamics of biological systems synthetically.

1.3. Cellular Scaffolding and Tissue Regeneration

There are two main uses for hydrogels in wound healing therapies, hydrogels which function as scaffolds to impart mechanical stability on the surrounding environment and promote tissue ingrowth, and hydrogels that serve as drug delivery depots which can be engineered to exhibit complex release kinetics for their associated therapeutic payloads. Although these two concepts are not mutually exclusive, and proper wound therapies will utilize both of these systems, the design constraints and engineering considerations that must be made during development differ substantially enough to merit their separate classification.

Hydrogel scaffolds are designed to provide a regenerative template or substrate onto which cells can adhere, proliferate and coordinate their responses to regenerate damaged tissues after injury (Fig 2b). These materials can either be delivered on their own²⁵ or along with pre-encapsulated cellular populations²⁶. In the former scenario, viable cells from the surrounding tissue migrate into the material from the periphery and interact within the scaffold to reconstitute the desired tissue at the site of implantation. The latter system proceeds largely through the same mechanism, with the important distinction that delivered cells can interact with native cell populations and deliver cytokines or other therapeutic secretions to recapture aberrant biology²⁷.

Figure 2. Key considerations for injectable hydrogel design.

(A) Over the past several decades numerous chemistries and material processing techniques have been utilized for the production of injectable hydrogels. These materials can be delivered either as liquid precursors which are crosslinked into stable gels through environmental triggers at the site of injection or as deformable solids which can withstand injection shear forces and re-anneal in the wound bed. In-situ annealing triggers include temperature, pH, light intensity, and concentration of ionic species and enzymes. (B) Hydrogel scaffolds function as regenerative templates that provide a suitable substrate for cellular ingrowth while matching the physiochemical properties of the native ECM. Matching scaffold stiffness to the surrounding tissue and optimizing its rate of degradation ensures that infiltrating cells remain viable, maintain their desired phenotype, and coordinate their response over the entirety of the wound healing process. Scaffolds can also be functionalized to provide sites for cellular binding or protein adsorption, enabling spatial control over the density of seeded cells as well as the availability of cytokines and growth factors. (C) Injectable hydrogels can also be designed to serve as depots for the controlled release of therapeutic compounds. Here, the release profile of compounds can be controlled through parameters such as polymer mesh size, polymer affinity for the target molecule, or rates of polymer degradation. Additionally, therapeutic compounds can be covalently linked to the polymer mesh by cleavable anchoring groups for controlled release upon exposure to external signals.

In general, these scaffolds take advantage of the structural similarity between the hydrophilic polymer meshes they are composed of and the natural extracellular substrates they are trying to replicate in order to aid regeneration wherever bulk tissue organization has been lost. Because the dynamics of cellular metabolism are regulated, to a large extent, by extra-cellular cues present in the local microenvironment, proper presentation and integration of these signals is critical to achieving the large scale cellular coordination necessary for tissue homeostasis. These extracellular factors manifest in the form of soluble signaling molecules, cell-cell interactions and external forces transmitted through the ECM via bound integrin receptors^28,29. One commonly utilized method of forming polymer-cell interactions is through the functionalization of polymer backbones with RGD peptides (arginine-glycine-aspartic acid) which are derived from domains of well characterized integrin ligands and promote cell attachment^30-32. RGD decoration enhances the efficiency with which regenerating tissues integrate with implanted biomaterials by providing localized sites of cellular integration, mediating foreign body inflammatory responses, and enhancing rates of vascularization. Newer systems allow patterning of designed materials with clustered RGD³³ and enable temporal control over peptide presentation through activation from external stimuli³⁴. Although these techniques will undoubtedly aid in the optimization of scaffolds by enabling the formation of cellular gradients throughout materials, they are limited by a reliance on a single well studied integrin binding sequence. Recent reports have demonstrated that preferential activation of specific integrin heterodimers have a tremendous impact on the resulting vascular morphology, with excessive activation of αVβ3 in particular, leading to endothelial sprout clumping and leaky blood vessels in vivo³⁵. Therefore, to truly take advantage of integrin binding peptides a more complete knowledge of the relationship between engagement and subsequent variation in cellular phenotype must be developed.

The degradation properties of hydrogels must also be considered whenever designing scaffolds for tissue regeneration. Because many injectable hydrogels form highly condensed structures with nanometer sized pores, micron scale proliferating cells are unable to infiltrate them without degrading the covalent bonds that hold them together. Thus, regeneration of wounded tissue requires the maintenance of a precise balance between the rates of tissue integration and scaffold degradation. Slow material degradation often results in increased inflammatory responses and can promote fibrosis³⁶. In contrast, materials that degrade too quickly provide insufficient scaffolding to maintain infiltration and bulk ordering of proliferating cells. To combat these limitations, several groups have reported methods to design injectable microporous scaffolds^37,38 which are able to accommodate tissue regeneration while retaining bulk stability. The widespread adoption of these systems will help alleviate design bottlenecks without sacrificing the tunability which allows scaffolds to meet the precise physiochemical requirements of the wound site.

1.4. Hydrogel Depots for Localized Drug Delivery

When treating diseases, drugs are often administered systemically, either orally through pills and tablets, or intravenously via injection. In many cases these delivery methods are suboptimal. Therapeutic molecules that are unstable in vivo due to susceptibility to enzymatic degradation or rapid clearance by the kidneys must be delivered at concentrations far higher than required for therapeutic effect³⁹. Additionally, the interactions of delivered molecules at off-target sites may result in unanticipated adverse effects, potentially complicating the health of patients even further⁴⁰.

Hydrogels functioning as depots for the localized release of therapeutic molecules serve as attractive alternatives to many of the current systemic approaches (Fig 2c). Manipulations in parameters such as pore size, backbone charge, hydrophilicity or cross-link density directly impact the diffusion rates and solubility of molecules dispersed throughout the hydrogel matrix. Thus, injected hydrogel depots can serve as sites of localized drug delivery over many time scales for a wide variety of therapeutic factors. Furthermore, novel on-demand drug release depots are being designed for controlled release upon application of external stimuli such as ultrasound, magnetic fields, electrical impulses or light waves⁴¹. These depots function by partially degrading when exposed to their associated stimuli and allow for tuning of release rates on a patient by patient basis after implantation. Self-healing on-demand depots have also been reported and add an additional level of control to drug delivery applications by re-annealing after the cessation of stimuli, ensuring that therapeutics are only released appreciably when a stimulus is applied⁴². Currently clinical use of hydrogel depots, is hindered by a reliance on intrinsic biological microenvironments for controlled release, where heterogeneity throughout the body and between individuals limits the efficacy of treatment approaches based on the location of implantation⁴³. Elimination of this dependency through the design of “smart” materials for on-demand release will be necessary to make hydrogel drug release depots more pervasive in the clinic for applications ranging from tissue regeneration, to treatment of complex diseases and immunizations.

2. Non-Covalent Hydrogels

We begin our examination of injectable therapeutic hydrogels with materials formed from reversible non-covalent crosslinks. These comprise many physically annealed hydrogels, which form due to changes in temperature or solution pH, but also includes ionically crosslinked gels, as well as hydrogels formed from self-assembling peptides (Fig 2a). In general, the hydrogels covered in this category can be formed fully outside of the body due to the additive effects of relatively weak internal attractive interactions such as hydrogen bonding, dipole-dipole forces, and hydrophobic interactions, are temporarily degraded through shear forces during injection, and readily reform in situ at the injection site.

2.1 Thermally Annealing Hydrogels

The necessity for sustained function in a diverse set of biological environments requires that precise control over the in situ formation of injectable hydrogel biomaterials be mediated by environmental cues that are constant both throughout the body and among patients. Temperature, is one such parameter, because it is both generally stable in vivo and can be readily controlled during ex vivo injection conditions. As a result, exploitation of temperature differentials to trigger sol-gel transitions of amphiphilic polymers has become commonplace in the field of biomaterial design since initial reports of the thermal properties of poly(N-isopropyl acrylamide) (PNIPAAM) first emerged in the 1960s⁴⁴. Nowadays, a vast library of thermally triggered polymers, including, PNIPAAM, Pluronics, poly(vinyl ether) (PVE), poly(N,N-diethylacrylamide) (PDEAM), poly(N-vinyl caprolactam) (PNVCa), and Poly(oligo(ethylene glycol) methyl ether metacrylate (PoEGMA) have been derived to form tissue engineering scaffolds, cell delivery vehicles, and drug release depots which omit the necessity of harsh reaction conditions found in many of their covalent counterparts⁹. Instead, these thermally annealed hydrogels, or thermogels, take advantage of a finely tuned balance between hydrophilic and hydrophobic groups of their backbone polymer to remain soluble in solution at room temperature, but collapse into micellular globules and self-associate at physiological temperatures^45,46 (Fig 3a). This transition, which occurs rapidly above the lower critical solution temperature (LCST), enables both rapid gelation upon injection, and uniform distribution of encapsulated cargo, making these materials excellent systems for a multitude of regenerative applications.

Figure 3. Overview of non-covalent gelation mechanisms.

Polymer backbones can be reversibly crosslinked into hydrogel networks through several methods that take advantage of interchain affinity. These crosslinking reactions include (A) self-organization of amphiphilic micelles at high temperature, (B) polymer bridging via multivalent ions, and (C) hierarchical self-assembly of short peptides. Source: Adapted with permission from Refs [13], [51], & [81].

Experimentally, thermogel drug release systems have been utilized for the sustained delivery of many molecules which exhibit particularly short half-lives in biological environments^47-49. Their increased protective capabilities stem from the polarity based self-association of polymer blocks and subsequent micellular reorganization which both isolates therapeutic payloads and limits interaction with external enzymes and ROS⁵⁰. Furthermore, release profiles are readily tuned through modifications to polymer molecular weight, polarity, and weight fraction resulting in easily modifiable delivery rates. Of note, Ci et al. reported high therapeutic efficacy towards the inhibition of colon cancer tumors in rodents through the sustained release of the topoisomerase inhibitor, irinotecan from poly(D,L-lactide-co-glycolide) (PLGA)/PEG triblock copolymers (PLGA-PEG-PLGA)⁵¹. Irinotecan, a well known anti-tumor drug suffered from high levels of inactivity when delivered through direct injection in its solubilized form due to a preferential to revert to its inactive carboxylate form at physiological pH. However, sustained delivery from PLGA-PEG-PLGA resulted in a five fold increase in the active lactone form of irinotecan throughout the entire two week release profile, alluding to the formation of a materials dependent bioprotective microenvironment. Similarly, Qiu et al. utilized PLGA-PEG-PLGA for the sustained delivery of platelet rich plasma (PRP) within full thickness rodent skin wounds⁵². Release of various growth factors from concentrated platelets allowed PRP to be utilized as an inexpensive cocktail of bioactive signals, resulting in more coordinated cellular response. In vitro, PLGA-PEG-PLGA PRP depots exhibited minimal cytotoxicity towards both fibroblasts and vascular endothelial cells and were able to maintain steady growth factor release over the course of two weeks, in stark contrast with un-encapsulated PRP which released over 90% of growth factors in the first 12 hours and had no detectable activity after four days.

Aside from serving as the driving force behind thermogelation, the amphiphilic nature of many thermogels also broadens the range of therapeutic payloads with which they are compatible. The solubility of hydrophobic small molecules, which is often not appreciable in aqueous environments can be substantially improved through preferential associations with hydrophobic domains of the thermogel. Variations in the degree of interaction between material and dissolved therapeutic lead to changes in release profile and can be honed to extend delivery up to a time period of months⁵³. As many FDA approved compounds are hydrophobic small molecules⁵⁴, amphiphilic thermogelling materials may offer a bridge to the application of many underutilized compounds within the field of regenerative medicine.

Efforts to utilize thermogels as regenerative scaffolds for improved wound healing outcomes have also demonstrated promise in rodent models. Injection of these materials into wound beds allows for rapid gelation in complex geometry, reinforcing the damaged tissue, and in some cases resulting in hydrogels with micron sized pores that allow for the diffusion of nutrients and signaling factors to improve phenotypic outcomes of seeded cell populations⁵⁵. Yun et al. examined the efficacy of one such thermogel in skin wound healing, by encapsulating fibroblasts within a 7.6 wt% PEG-poly(L-alanine) (PEG-PLA) hydrogel⁵⁶. This material formed a relatively strong (600 Pa storage modulus) hydrogel and was able to host encapsulated fibroblasts, which exhibited a spherical morphology and aggregated to form 30-80 micron sized clusters. Cell proliferation as measured by CCK-8 demonstrated a tenfold increase in cell number over the course of three weeks alluding to the cytocompatibility of this material. Subsequent evaluation of collagen mRNA expression demonstrated elevated collagen III production at early time points as well as increased collagen I production throughout the study in comparison to fibroblasts cultured in a matrigel control. This response is more physiologically relevant, as formation of new tissues usually progresses through an initial deposition of a provisional ECM rich in collagen III which is subsequently replaced by a more mature collagen I dense matrix⁵⁷. Lastly, comparison of the fibroblast seeded PEG-PLA with naked PEG-PLA and phosphate buffered saline (PBS) controls demonstrated a significantly increased rate of wound closure with formation of epithelium and granulation tissue at day 14 and fully filled epidermis decorated with hair follicles and sebaceous glands by day 21 in fibroblast PEG-PLA only.

Several studies have also focused on soft tissue regeneration with thermogels containing adipose derived stem cells. Tan et al. demonstrated promising in vitro cell viability and migratory effects within NIPAAM grafted HA gels but observed limited infiltration in vivo five days after subcutaneous dorsal injection in mice⁵⁸. More recently, Kim et al. reported an adipose derived soluble ECM (sECM)/methylcellulose hybrid thermogel seeded with human adipose derived stem cells (hASCs) for the treatment of skin wounds⁵⁹. Methylcellulose on its own displays thermogelling properties at elevated temperatures (40°C-60°C) stemming from hydrophobic interactions, but synergistic interactions with anionic groups on sECM was able to significantly lower the LCST for the composite material to promote gelation at body temperature. Live dead assays on seeded hASCs demonstrated material cytocompatibility in 6 wt% gels and subsequent material injection into full thickness rat wounds showed both uniform hASC distribution and significant infiltration by surrounding cell populations after 3 days. Thermogel cell depots again exhibited increased rates of wound closures relative to controls and resulted in minimal scar formation.

Although these studies provide reason to be optimistic there are limitations which inhibit purely thermogelling materials from optimal behavior. Secondary effects such as variations in the pH or oxidative stress of the wound bed can lead to variability in therapeutic outcomes and should also be controlled for optimal biomaterial application^60,61. Additionally, temperature gradients, especially for materials interfacing with the outer environment, may reduce tissue incorporation and physical properties across the treatment site. Therefore, thermogelling systems should not be considered a complete solution to impaired regenerative responses, but more of a stepping stone to a better understanding of material properties for improved therapeutic design.

2.2 pH Dependent Gelation

A second mechanism with which hydrogel gelation can be induced in situ is via pH mediated crosslinking. Here, the susceptibility to ionization of various pendant functional moieties is harnessed to generate attractive or repulsive forces between different segments of polymer backbone, resulting in both reversible dynamic crosslinks and controllable release kinetics for drug delivery applications⁶². In chronic wounds, the local pH is highly variable and has been measured as low as 5.4^63,64 and as high as 8.9⁶⁴ depending on the location of the wound, degree of necrosis and local oxygen availability. This deviation enables pH inducible hydrogelation to be utilized as a potential treatment modality for skin associated malignancies.

The library of polymers used to generate pH responsive hydrogels is broad. The ubiquity with which functional groups can be modified through oxidation or reduction implies that most polyacids, polybases and naturally occurring polymers can be utilized to form pH responsive hydrogels. Irrespective of the polymer backbone identity, charge-charge interactions between polymer and drug form favorable domains within these material that have proven advantageous for the modulation of release kinetics. Polymers that are tunable through separate mechanisms, such as the thermogels discussed above, can also be modified with pH responsive groups to provide an additional tuning parameter and may be used to trigger instantaneous transitions in shape, opening the door for pH actuated, smart hydrogel drug delivery⁶⁵. The relatively weak interchain interactions of these hydrogels unfortunately limits their utility as tissue scaffolds. Thus, the creation of stable and injectable cell culture environments is often reliant on secondary reinforcing interactions which are either ionic⁶⁶ or covalent⁶⁷ in nature. Nevertheless, the buffer like properties provided through the hydrogen ion absorption effects of these materials are extremely useful for mitigating excessive deviations from homeostasis in chronic wound environments and are being broadly applied in many regenerative systems. Recently, Koehler et al. reported on the formation of a pH sensitive IPN formed from PEG-diacrylate (PEGDA), acrylic acid and alginate which could be used to improve healing outcomes in chronic injuries⁶⁸. This material demonstrated high liquid uptake of up to 500% of the initial hydrogel mass at equilibrium, high buffering capacity under basic conditions and increased cell migration rates in both 2D and 3D in vitro healing models. The optimization of such properties would allow for removal of excess wound exudate, reduction in pH deviation to recapture normal non-malignant biology, and rapid tissue closure, all desirable properties for the treatment of slow to heal wounds. Therefore, in spite of the reduced mechanical properties provided from these interactions, newly designed injectable hydrogels will need to be carefully tuned in their pH responsiveness to properly control and reverse chronicity.

2.3 Ionically Crosslinked Materials

Ionic crosslinks provide yet another method of reversible bond formation within hydrogel biomaterials. These systems take advantage of interactions between localized charges on the surface of polymer backbones and counterions present in solution to form salt bridges that intertwine normally linear macromolecules to form mechanically stable gels (Fig 3b). In general, ionic crosslinking agents can be divided into two classes, multivalent ions, and charge bearing small molecules⁶⁹. Although the bonds formed by either of these classes of crosslinking agents are identical, their application can result in vastly different outcomes due to variation in diffusion rates, shielding effects of secondary ions in solution, and the dependence on pH and pKa of both the backbone polymer and small molecule cross linkers.

In general, the crosslink density of a hydrogel is an important parameter because it is easily manipulated and directly affects mechanical strength, material swelling, and diffusion rates of loaded therapeutic biomolecules. However, to date, precise control over the gelation kinetics of ionic systems has been difficult to achieve, and commonly used methods of gel formation such as dropwise addition of polymers to ionic solution result in highly heterogeneous materials with reduced mechanical stability, unconstrained degradation and limited cellular infiltration⁷⁰. Numerous efforts to control these parameters have been reported but often result in limited success. For example, Joddar et al. attempted to increase the strength and stiffness of CaCl₂ crosslinked alginate hydrogels through the incorporation of carboxylic acid functionalized multi-walled carbon nanotubes (MWCNT)⁷¹. Although the MWCNT alginate gels showed promising results as a culture medium through increased ECM deposition and migration of seeded HeLa cells relative to alginate controls, observed variations in physical properties did not follow a well-defined trend and unexpectedly became more brittle when mixed with certain ratios of MWCNT. Similarly, efforts to control the release profiles of encapsulated vascular endothelial growth factor (VEGF) by varying the ionic ratio used to form alginate microparticles showed preliminary feasibility but alluded to numerous difficulties such as a preference for ions to bind specific monomer pairs, differences in the cytotoxicity of ionic cross linkers and an inability to vary release profiles by simply mixing ionic species in the crosslinking solution⁷². Therefore, there is still considerable room for improvement in terms of expanding the tunability and reproducibility of these ionic systems.

Even with the difficulties in maintaining homogeneity, ionically crosslinked hydrogels have demonstrated preliminary success when utilized as therapeutic depots, both for the treatment of antibiotic resistant biofilms⁷³ and the rapid regeneration or revascularization of damaged tissues⁷⁴. Notably, Schmitt et al. demonstrated that ionically crosslinked hybrid alginate/0.1% PEG/0.1% HA gels could be used as injectable carriers for the delivery of autologous mesenchymal stem cells (MSCs) into wound sites post injury⁷⁵. Upon optimization of gel formation this system showed numerous desirable properties including a gelation time of under 30 minutes, adhesive tendencies towards an underlying collagen substrate, injectability through a standard needle and the ability to maintain viable MSCs for up to 6 weeks in vitro. Subsequent analysis of cellular growth factor delivery from these scaffolds demonstrated that both VEGF and basic fibroblast growth factor (BFGF) continued to be released at appreciable levels over the entire 6 week duration of the study. Because autologous stem cells are easily harvested and generate numerous vital growth factors on their own, optimized delivery vehicles have the potential to drastically improve regenerative outcomes without the exorbitant cost associated with repeated injections of intrinsically unstable recombinant proteins and therapeutic molecules or the safety concerns of gene transfer.

Ionic scaffolds have also been utilized extensively to probe the link between material properties and cellular phenotype. It is now readily accepted that the mechanics of substrates impact the phenotype of cells cultured on them, and may induce changes in proliferation⁷⁶, motility⁷⁷ or differentiation⁷⁸. For example, stem cells grown on stiffer substrates exhibit a tendency to preferentially differentiate into osteogenic lineages, whereas moderate elasticity leads to myogenic differentiation and softer substrates encourage differentiation towards neuronal or adipogenic cell types. Recently, efforts to elucidate the underlying pathways that govern these material-cell interactions have uncovered the importance of often overlooked mechanical parameters. In particular, stress-relaxation, or the ability for a material to dissipate applied forces and relieve stresses over time has proved significant. Chaudhuri et al. were able to synthesize a range of alginate based ionic hydrogels with identical initial elastic moduli but different rates of stress relaxation (Fig 4a-b). Interestingly, they observed that stem cells grown on these substrates differentiated with varying efficiency⁷⁹. In short, although substrate stiffness was still the predominant determinant of differentiation lineage, adipogenic differentiation and osteogenic differentiation were significantly increased on soft scaffolds with low rates of stress relaxation and stiff scaffolds with high rates of stress relaxation respectively (Fig 4c). Although the mechanism underlying this effect is not fully understood researchers have demonstrated that the ability for cells to rearrange their surrounding environment plays a role. This scaffold rearrangement allows cells to cluster integrin binding ligands to activate intracellular signaling pathways and to spread and form complex networks that aide intercellular communication. As the bonds formed through ionic crosslinking are reversible, shear thinning behavior is an intrinsic behavior of these systems⁸⁰. Numerous methods of tuning stress-relaxation in ionic systems have been reported, including through changes to the molecular weight of the backbone polymer, by utilizing a mixture of both covalent and ionic bonds⁸¹, and by incorporation of steric spacers. As the interplay between scaffold rearrangement and cellular phenotype becomes more understood, ionically crosslinked materials may prove particularly useful in the delivery of stem cells for regenerative applications.

Figure 4. Material rearrangement can induce changes to cellular phenotype.

(A) Chaudhuri et al. synthesized ionically linked alginate hydrogels with identical elastic moduli but varying rates of stress relaxation by simultaneously modulating the molecular weight of crosslinked alginate and incorporating PEG spacers into the polymer backbone. (B) Fibroblasts encapsulated in rapidly stress relaxing hydrogels exhibited increased spreading due to their ability to dynamically rearrange the surrounding scaffold. (C) Interestingly, the rate of stress relaxation also influenced the efficiency with which cultured mesenchymal stem cells differentiated into different lineages. Here, stem cells grown on stiff substrates with high rates of stress relaxation had a tendency to mature towards an osteogenic phenotype, whereas stem cells grown on slowly relaxing soft substrates efficiently differentiated towards a more adipogenic phenotype. (D) Stress relaxing materials have also been designed using other chemistries. In one example Rodell et al. combined guest-host complexation with covalent crosslinking to form stiff stress relaxing hydrogels which showed promise as regenerative scaffolds for tissue engineering applications. The routine consideration of stress relaxation rates in hydrogel scaffolds may aid the design of new materials which provide a greater degree of control and uniformity in response for both pre-seeded and infiltrating cell populations. Source: Adapted with permission from Refs: [79] & [135].

2.4 Peptide Self Assembly

Although polymeric biomaterials are widely used as tissue engineering constructs in the biomedical fields they are often plagued with issues such as the presence of undefined concentrations of impurities and batch-to-batch variation in composition within synthesized material products. Unexpected deviations in these properties may present confounding effects which limit the reproducibility of studies reliant upon their interactions with cultured cell populations. Self-assembling peptide based hydrogels comprise a class of materials that are commonly used as alternatives to polymer based systems.

Derived from naturally occurring protein structural motifs, self-assembling peptides are short chains of amino acids with alternating domains of charge and polarity. When dissolved in neutral solvents and physiological salt concentrations, these peptides interact via hydrogen bonding, ionic bonding, hydrophobic interactions or van der waals forces to spontaneously assemble into hierarchical nano-scale structures such as vesicles, particles, tubes and fibers¹² (Fig 3c). Materials derived from these assemblies have the additional advantage of being non-toxic, non-immunogenic, non-thrombogenic, degradable, and easily metabolized as the degradation products are naturally occurring amino acids⁸². Of all the peptide nano-assemblies, nano-fibers are particularly useful for the design of tissue regenerative hydrogels. These fibers are on the same size scale as natural ECM fibers⁸³, can be easily designed to mimic the stiffness of a wide range of soft tissues⁸⁴, and can be further functionalized with the attachment of cellular interacting peptide domains or cytokines and growth factors⁸⁵, allowing for the design of biologically relevant culture environments and improved control of proliferating cellular populations. For example, Bradshaw et al. report that functionalization of widely studied RADA16 peptides (RADARADARADARADA) with a fibronectin attachment motif (RADA16-GG-RGDS) or a collagen type-1 derived motif (RADA16-GG-FPGERGVEGPGP) leads to significant changes in both proliferation and migration of cultured fibroblasts and keratinocytes relative to unmodified controls⁸⁶. These findings indicate that spatial gradients in self-assembling peptide materials have the potential to induce diverse cell responses at various locations within a regenerative scaffold.

Self-assembling peptides have also been utilized for the promotion of angiogenesis within regenerating tissues. Because supply of nutrients and dissolved gases at levels sufficient to promote long term survival of tissues is limited to a distance of 200 microns from the supplying vessel, sufficient vessel formation throughout reforming tissues has limited the efficacy of many therapeutic systems⁸⁷. Wang et al. utilized a mixture of similar RADA16-I fibers functionalized with either a PRG peptide containing two RGD binding sequences or a previously engineered peptide KLTWQELYQLKYKGI (KLT) (which interacts with VEGF receptor to induce activation of the VEGF dependent proliferation pathway) to form RADA hydrogels⁸⁸. As expected RADA16-PRG fibers increased binding of human umbilical vein endothelial cells (HUVECs) at initial time points and RADA-16-KLT promoted the formation of capillary like structures through the arrangement of cultured endothelial cells after 1 day. Of note, these RADA based hydrogels were able to induce migration of endothelial cells from non-functionalized peptide regions, organize these cells into structures with vascular morphology, and sustain the developed vasculature without the use of any exogenously delivered growth factors alluding to the importance of permissive scaffolds for proper endothelial maturation.

Efforts to utilize peptide based scaffolds for the regeneration of partial and full thickness skin wounds have also been successful both in vitro and in vivo. Schneider et al. utilized RADA fiber hydrogels for sustained delivery of epidermal growth factor (EGF), a molecule with a relatively short half-life to induce proliferation and migration of keratinocytes in a human skin equivalent (HSE) model⁸⁹. Release from this depot was completely controlled by hydrogel degradation through proteolytic enzymes in the wound bed as observed through in vitro release assays, where 65.5% of total EGF was released after 48 hours in the wound bed, versus no measurable release in PBS over the same time frame. Regeneration rates of HSEs following incisional wounds were dramatically increased when treated with EGF gels showing ∼60% closure after 48 hours in contrast to ∼14% for control. In vivo, Meng et al. used RADA nano-fiber gels to treat a partial-thickness burn model in sprague-dawley rats⁹⁰. Burns treated with RADA increased contraction at all time points in comparison to chitosan, poly-(DL)-lactic acid, collagen and sham wounds, and were unique in that they prevented formation of blisters at the wound site and halted early time point expansion of the wound area due to edema.

Although the RADA based nanofibers we have focused on above are among the most widely used self-assembling peptide systems for injectable hydrogel design, they are by no means the only option available. Simple short peptide derivatives such as fluorenylmethoxycarbonyl (Fmoc) have also been assembled into nanofibers⁹¹ and decorated with RGD peptides⁹² to form biocompatible scaffolds which uniquely present cell receptor ligands from fiber surfaces to more accurately mimic important ECM interactions. The use of self-assembling peptides is a relatively recent development within the realm of injectable therapeutic hydrogels, hence their application is still in its infancy. As, we learn more about the structure-function relationships of proteins, increasingly advanced self-associating peptides will be engineered to generate a variety of structural motifs. Recent studies have demonstrated that recombinant protein structures⁹³ or non-native amino acids⁹⁴ can be engineered to reinforce self-assembling materials, conjugate to therapeutic payloads, and elicit novel biological responses. These newer synthetic protein materials will aid in the development of more complex and efficacious hydrogels for regenerative medicine.

3. Covalent Hydrogels

Covalently bound hydrogels have also found widespread use in the field of regenerative medicine. Unlike the bonds which compose physically annealed hydrogels, covalent crosslinks are not inherently reversible and therefore usually result in sturdier, more mechanically stable materials. This increased stability allows covalent hydrogels to serve as more mechanically robust scaffolds than purely physical hydrogels and enables the physical matching of a wider array of tissue types. In the past, delivery of these materials relied upon highly invasive surgical procedures, but newer mechanisms of formation either through photo induced radical polymerization or delayed bonding via reactions with slow kinetics have enabled transfer to the site of interest through non-invasive injection (Fig. 2a). Once delivered, the stronger binding forces prevent covalent gels from diluting with the surrounding fluids or diffusing from the injection site, leading to efficient and long-lasting incorporation of the material within native tissues. In this section and the one to follow, we hope to highlight some of the important properties of covalently formed hydrogels, first through light induced gelation approaches and subsequently through two component delayed crosslinking chemistries.

3.1 Photocrosslinking

Light induced hydrogel polymerization (Fig 5a) proceeds through a three-step radical chain growth reaction mechanism¹¹. Initially, a solution of small molecule photoinitiators, and monomers is illuminated at specific wavelengths in order to excite and cleave specialized bonds within the photoinitiator molecules. This process leads to the creation of free radicals which subsequently attack and withdraw electrons from double bonds present within the backbone monomers and produce new radicals at the opposite ends of the attacked bonds. The new monomer radicals then progress through a period of propagation wherein they attack similar unsaturated groups in neighboring monomers and elongate into a macromolecular polymer chain. During the early stages of propagation, the concentration of unreacted monomers far exceeds the concentration of activated free radicals and the polymerization rate rapidly accelerates until reaching a maximum. Finally, however, the concentration of radical chains becomes significant and interactions between pairs of radical groups begins to dominate, leading to the quenching of both radicals and termination of polymer growth.

Figure 5. Application of light improves spatial control of hydrogel formation and drug release.

(A) Polymers decorated with radical generating functional groups can form hydrogels in the presence of photoinitiators and appropriate wavelengths of light. (B,C) When combined with specialized projection techniques or photomasks these systems can be utilized to form complex 3-dimensional geometries or induce localized binding and release of therapeutic payloads. (D) De Forest and coworkers demonstrated the utility of such light mediated controlled release schemes by utilizing orthogonal photoactivated chemistries for the binding and subsequent release of full length proteins. Here hydrogels were pre-formed and decorated with photocaged alkoxyamine groups. When illuminated with UV light through patterned photomasks, photocages were selectively degraded. Subsequent addition of aldehyde functionalized proteins allowed spatial patterning of bioactive molecules. Introduction of o-nitrobenzyl esters into protein linkers allowed for subsequent photorelease upon exposure to similar UV irradiation. (E,F) To highlight this technology, multiphoton laser scanning lithography and photolithography demonstrated spatial control over photorelease of fluorescently tagged proteins in 3D and 2D respectively. Source: Adapted with permissions from Refs [103], [104], [107], & [110].

The benefits of photopolymerization are numerous. The rapid and specific nature of the reactions provides improved spatio-temporal control through the location and duration of applied light⁹⁵. Heat generated over the course of polymerization is readily constrained through changes in light intensity, photoinitiator properties, or the application of light in intervals over the course of gelation⁹⁶ which allows material compatibility with both cells and biomolecules that may otherwise be rendered inactive due to local temperature extrema. The choice of photoinitiator, may also lead to variation in the final material properties and cytocompatibility of the hydrogel, although this design parameter is not as thoroughly characterized. A preliminary study by Mironi-Harpaz et al. examined how variations in photoinitiator type, concentration, and light intensity affected the shear modulus, reaction rate, and cytotoxicity towards encapsulated cells of photogelling materials⁹⁷. Although the authors were able to differentiate between several photoinitators, with Igracure-2959 and Igracure-184 performing better than their counterpart in both kinetic rate and cell viability assays, this study was limited in scope, and more data will need to be collected before photoinitiators can be selected for optimal performance a priori. Even so, several other reports have demonstrated that at low photoinitiator concentrations, photocrosslinking allows for encapsulation of many different cell types with minimal cytotoxicity^98,99, but may result in altered gene expression due to the high prevalence of free radicals ¹⁰⁰ implying that these chemistries are compatible with regenerative applications but should be assessed carefully prior to use as depots for cellular delivery. Finally, when utilized for in vivo applications, photogelling polymers can be injected as liquid precursors at the site of interest and cured in situ for a rapid and minimally invasive implantation. Even with these benefits however, it should be noted that many of the photoinitators currently relied upon are activated by light in the UV range, which can penetrate tissues deeply and damage the body. Widespread adoption of these materials for clinical wound healing and drug delivery applications will necessitate the design of efficient photoinitiator molecules which can be triggered at higher wavelength visible or infrared light sources.

Recently there has been much focus on combining light induced hydrogel polymerization with innovative physical or chemical systems to design physiologically relevant and heterogeneous tissue culture environments. Standard photolithography has been widely adopted to polymerize 2D patterns through a photomask, enabling both the design of surface microarchitecture and establishing gradients of bioactive peptide sequences within the material substrate¹⁰¹. Only recently, have these technologies been extended to tackle the nontrivial task of patterning biomaterials in 3D to allow refined examination of small molecule diffusion effects, and cell-cell or cell-material interactions. Of note, Soman et al. were able to use a previously reported dynamic projection stereolithography system¹⁰² for the formation of complex three dimensional shapes within natural polymer hydrogels containing encapsulated populations of cells¹⁰³ (Fig 5b). In this method, a digital mirror device is used to dynamically project a UV source onto a macromer solution in a well. The patterning of this projection is dictated by user defined CAD files, and are changeable over time. By dynamically shifting the polymer precursor well upwards as the mask is changed, complex 3D geometries can be formed within the polymerizing cell-seeded material and cellular dynamics can be studied within a wide variety of predefined shapes.

Although these methods of geometrical modifications are interesting they are not easily applied to materials that will be injected into a patient post processing. Perhaps more relevant to actual injectable treatments are approaches to dynamically and reversibly attach or release small molecules and proteins from the scaffold itself¹⁰⁴ (Fig 5c). For example, degradative approaches can be used for the on demand release of therapeutic molecules from localized depots or to locally rearrange hydrogel properties to direct cellular growth and migration¹⁰⁵. Other approaches have been reported to allow the widespread functionalization of hydrogels with full length proteins that would otherwise be denatured from high intensity light exposure. For example, Mosiewicz et al. embedded proteins protected by a photodegradable cage within a hydrogel functionalized with a counter-reactive substrate in the presence of activated factor XIII¹⁰⁶. Under standard conditions this reactive system remained inert, but upon exposed to UV radiation the protective cage was degraded and the protein of interest was bound to the hydrogel through a factor XIII mediated coupling reaction. Although the scope of this study was limited to a few commonly used growth factors it can in theory be repurposed for the decoration of any desired protein within designed biomaterials. Extending the concept of protein patterning even further, studies by De Forest et al. resulted in a clever photochemical method of spatially controlling the ligation and release of proteins through UV exposure of a photocaged alkoxyamine hydrogel in the presence of o-nitrobenzyl ester modified proteins¹⁰⁷ (Fig 5d-f). This reaction scheme allows for the three-dimensional patterning of multiple proteins at predictable concentrations and with micrometer level precision. The authors allude to the utility of this material in assaying biomolecule presentation in cell culture, but one can also envision how the knowledge obtained from these studies could inform the development of optimal protein ratios or release profiles for in vivo application.

Efforts to utilize light induced hydrogel photopolymerization for epidermal regeneration and wound healing have also resulted in modest successes, but have for the most part been limited to a few commonly used natural polymer systems. In particular, work from Zhao et al. has demonstrated the efficacy of methacrylated gelatin (GelMA) based materials towards epidermal reconstruction^108,109. These GelMA hydrogels provide significant variation in both mechanical properties and material stability through simple gelatin concentration variations in the pre-polymer solution. By optimizing these parameters, the authors were able to develop GelMA hydrogels that had similar moduli to natural epidermis, maintained the growth, proliferation and self-organization of keratinocytes, and degraded in a timespan of months, making them ideal candidates for regenerative scaffolds in chronic wounds. Follow up studies, showed that these materials were compatible with widely used material processing techniques, such as electrospinning, and resulted in increased rates of wound closure in in vivo murine wound healing assays. Supporting work also showed that GelMA/HA-MA hydrogels loaded with adipose derived stem cells allow high levels of proliferation and vascularization, further highlighting their utility in wound healing¹¹⁰. However, widespread utilization of these materials in the clinic is again limited by the need for high intensity UV light with many photoinitator molecules, and new photoinitators will have to be designed before photogels can be used for many applications in vivo.

3.2. Injectable Covalent Precursors

The design of covalently linked hydrogels is not strictly dependent on the use of radically propagated photoinitiation reaction schemes. Rather, there are a host of biocompatible chemistries whose kinetics are slow enough to enable mixing ex-vivo, injection as a liquid precursor solution into the site of interest, and final solidification and annealing with the surrounding tissue environment. Of these systems, some of the best known reactions include acrylate homopolymerization¹¹¹, Michael type addition¹¹², oxime linkage¹¹³, and Schiff base polymerization¹¹⁴. For chemistries that require the use of free radicals, redox approaches can be utilized as a substitute for light with molecules such as ammonium persulfate (APS) and tetramethylethylenediamine (TEMED). Over the past decade, there has been much interest in a subset of these covalent precursor reactions, termed “click” reactions, within the field of material science. These click chemistries were designed on the premise that one could achieve a tremendous amount of synthetic mileage with relatively few reactions, provided that they were modular, orthogonal, proceeded under mild reaction conditions, had high yields of a single product, and required few back-end purification steps¹¹⁵. Common click reactions include strain promoted azide-alkyne cycloaddition¹¹⁶, diels-alder cycloaddition¹¹⁷, and nitrile oxide-norbornene cycloadditions¹¹⁸, in addition to the Michael type thiol-ene additions¹¹², and oxime linkages¹¹³ mentioned above. Initially, several of the reported click reactions relied upon the use of copper catalysts. However, these soon proved to be relatively cytotoxic and thus not compatible with applications in the presence of cells or tissues. This has prompted a whole branch of research into the development of copper free click reactions, to allow for biorthogonality and biocompatibility¹¹⁹. Regardless of the application, this library of chemistries provides the versatility and ubiquity needed for many drug delivery or wound healing applications.

Delivery of small molecules from covalent gels is straightforward and can be accomplished by simply mixing the active ingredients with the hydrogel precursors prior to injection. This results in a uniform distribution of drug within the delivered polymer and is only limited by the solubility of the therapeutic ingredient. Zhang and colleagues, recently applied this technique to interrogate the importance of the master regulator HIF-1α within the wound healing cycle¹²⁰ (Fig 6). HIF-1α is one half of a hypoxically induced protein enhancer complex which plays a role in tuning the expression levels of hundreds of genes¹²¹ and was shown to be heavily under expressed in healing impaired strains of mice. In order to prolong the half-life of these proteins in vivo an 8-arm PEG cysteine/N-hydroxysuccinimde based hydrogel¹²² was formed for the sustained release of 1,4-dihydrophenonthrolin-4-one-3-carboxylic acid (1,4-DPCA), a small molecule which inhibits the activity of degradative prolyl hydroxylases¹²³ (Fig 6a). In vitro fibroblast culture with these gels indeed resulted in overexpression of HIF-1α. When injected into the back of the neck of regeneration inhibited, Swiss Webster mice HIF-1α expression was again recovered, and could be detected at high levels for a period of five days. The authors next wounded the ears of Swiss Webster mice and treated them with three hydrogel injections at five day intervals. Complete healing of these wounds was observed by day 35 in treated mice only (Fig 6c-d). Furthermore, healing responses were readily reverted by delivery of HIF-1α specific siRNA, solidifying the importance of HIF-1α in tissue regeneration and demonstrating the potential utility of 1,4-DPCA loaded hydrogels as regenerative materials for hard to heal wounds. Before this technology can be adapted clinically, further studies will have to be done to evaluate the significance of HIF-1α in the regenerative response of other mammals as well as in other tissues.

Figure 6. Sustained release of drugs from covalently bound hydrogel networks can aid wound repair.

(A) Zhang and coworkers encapsulated 1,4-DPCA, a potent inhibitor of prolyl hydroxylases into an injectable hydrogel for sustained release in vivo. Inhibition of prolyl hydroxylases stabilizes HIF-1α, a central molecule in many regenerative processes, and allows for examination of its contribution to the wound healing process. (B) The drug eluted continuously from the material over the course of several days. (C) Mice were ear punched and treated with either gel alone (G₀) or gel loaded with 2 mg/mL 1,4 DPCA (G_d). At day 35, after three successive injections of gel, wound healing was visibly improved for G_d treated wounds over G₀ controls. (D) Histological analysis of regenerating ear at day 35 showed more complete wound closure for G_d treated wounds along with a greater number of mesenchymal cells in the regenerating bridge when compared with G₀ controls. Further analysis (not shown) demonstrated that HIF-1α blockage using siRNA reverted this regenerative improvement, supporting the hypothesis that upregulation of HIF-1α is sufficient to improve regenerative outcomes. Source: Adapted with permission from Ref [120].

Covalent precursors are also widely used to form injectable scaffolds. For example, Li and colleagues took advantage of Schiff-base crosslinking to form a scaffold from carboxymethyl chitosan and oxidized dextran that showed good cytocompatibility in vitro and enhanced regeneration rates in vivo within rats with second degree burn wounds¹²⁴. Through judicious choices in polymer backbone chemistry researchers have also demonstrated effectiveness in sequestering pro-inflammatory cytokines within scaffolds¹²⁵, leaving the door open for the development of materials which can promote infiltration while simultaneously mitigating undesirable chemical signals in chronic wounds. In this light, Feng et al. reported on the formation of a combination Konjac glucomannan/heparin enzyme crosslinked scaffold for the localized activation of macrophages and sequestration of their released pro-angiogenic signals¹²⁶. These gels formed in a matter of minutes after mixing with their activating enzyme, and preliminary evaluations in vitro confirmed their ability to bind proangiogenic growth factors VEGF, platelet derived growth factor-BB (PDGF-BB), and BFGF. Both suspended human monocytes and anchorage dependent murine machrophages adhered to this material rapidly and demonstrated upregulated production in proangiogenic signals. In vivo these gels again displayed higher levels of VEGF and PDGF sequestration than gels which were not crosslinked with heparin, and showed widely distributed vasculature after two weeks, a phenomenon that was absent from all control gels. By optimizing the design of such systems hydrogels have the potential to localize signals from the patient's own wound environment to promote therapeutic outcomes without the need for exogenous factors.

4. Hybrids

Alongside the vast library of physically, ionically, and covalently crosslinked hydrogels described above, there are several hybrid design schemes, which are not as easily delineated into these categories. These hybrid hydrogels may combine traits from multiple crosslinking classes, or may be unique in that the injection of fully crosslinked gels or gel building blocks may be feasible. In the sections to follow we choose to highlight three of these hybrid systems, dual crosslinked gels formed by guest-host interactions and reinforced with covalent bonds, cryogels, and microporous annealed particle hydrogels (Fig 2a).

4.1 Physical-Covalent Hybrids

The use of synthetic hydrogels as regenerative biomaterials has often fallen short in clinical outcomes due to an inability to recapitulate the complexities inherently present within the natural ECM. Far from a static scaffold, the ECM provides a reciprocally dynamic microenvironment that both regulates cellular behavior through the release and sequestration of chemical signaling molecules and that is regulated through cellular mediated degradation and subsequent reorganization in response to physiological changes¹²⁷. In contrast, many scaffolds are formed from relatively rigid non-biodegradable covalent crosslinks, resulting in almost purely elastic behavior. Although these systems provide the structural stability desired in regenerative applications, they often impede many desired cellular functions such as migration, proliferation and phenotypic transformations^79,128,129. In order to circumvent these limitations, physically associating guest-host interactions have been paired with common covalent cross-linking schemes to form dual physical-chemical hydrogels (Fig 4d).

Guest-host interactions form from the preferential association of polymer backbones functionalized with macrocyclic cavitand hosts, such as cyclodextrins¹³⁰, crown ethers¹³¹, cucurbit[n]urils¹³², calix[n]arenes¹³³, and pillar[n]arenes¹³⁴, with smaller guest molecules which both match the size of the macrocyclic pocket and induce compatible electronic interactions. When thoroughly mixed these backbones reversibly self-assemble into hydrogel matrices which are both injectable and highly malleable under the forces exerted by infiltrating tissues. Moreover, the interactions between guest and host enable higher ordered structural assemblies to form within functionalized polymers simply by the progression towards thermodynamic equilibrium, and endow these hydrogels with self-healing properties. Finally, the proportion of physical guest-host interactions to covalent cross links is easily varied, and can be utilized to match tissue mechanical properties while preventing material dissipation from the site of injection¹³⁵. Presently the use of guest-host hydrogels is limited to mostly in vitro applications¹³⁶. However, these materials carry tremendous potential for tissue regeneration, and may theoretically enable the development of a class of injectable scaffolds which can achieve a wide range of material properties while remaining deformable on a cellular scale.

Therapeutic delivery of molecules from guest-host materials is aided by the formation of inclusion complexes between host and payload. The quasi specificity of guest-host interactions implies that dissolved small molecules can also localize to the macrocyclic pocket, provided they have the right size and polarity to induce attractive interactions. This can both increase the solubility¹³⁷ and prolong the availability of unstable drugs for in vivo applications¹³⁸. These advantages can be exploited in physical-covalent hybrid hydrogels by utilizing host functional groups as both the crosslinking moiety and site of drug loading¹³⁹. For example, Mateen and coworkers reported on a β-cyclodextrin (β-Cd/dextran hydrogel crosslinked through hydrazine bonds for the sustained release of dexamethasone¹⁴⁰. In this work the authors achieved an over 30-fold increase in dexamethasone loading simply through the addition of 11 wt% β-Cd over control gel, with only ∼5% release over the course of a 20 day observation time frame. Of note, release profiles from these systems are no longer strictly governed by molecular diffusion through the porous hydrogel medium, but are also impacted by the binding constant between drug and host. Therefore, modulation of the functional group ratio in traditional guest-host gels impacts release rates much more so than in covalent guest-host hybrids, where the binding pocket is unperturbed during crosslink formation¹⁴¹.

Exploration of the effectiveness of guest-host hybrids functioning as tissue engineering scaffolds is limited due to the general lack of in vivo data from these systems. However, several previously reported studies offer a glimpse that their use towards this end is at least feasible. One interesting proof of concept study from Highley et al. highlighted the compatibility of guest-host hydrogels with 3D printing technology¹⁴². Here, HA based hydrogels functionalized with methacrylate, cyclodextrin, and adamantane were printed in complex 3D geometries within a support gel and subsequently exposed to UV radiation to induce formation of secondary covalent bonds. This process enables the formation of cell culture environments with physiologically relevant geometries, such as mock vasculature, which can be exposed to pressure driven shear flow, and provides a direct platform to establish physio-chemical gradients within the hydrogel itself. Additional works have also demonstrated the efficacy of these hybrid materials towards retention and incorporation within natural tissues¹³⁵, as well as towards the infiltration and maintenance of native cellular populations¹⁴³. Currently clinical adoption is limited due to incomplete knowledge of the degradation behavior, immune response, and toxicity profiles ascribed to these materials within the context of a biological host. However, with continued progress, it is only a matter of time before guest-host hybrid gels are widely adopted for use in in vivo studies and regenerative wound healing applications.

4.2 Cryogels

The materials discussed thus far have been injectable largely due to their liquid-like behavior under ex-vivo delivery conditions. In situ assemblies of this form are inherently adaptable and can be applied to many different biological environments through clever choices of polymer backbone, functional side chains and cross-linking chemistries. Unfortunately, the similarity in application between many of these systems also leaves them with the same limitations, namely leakage into surrounding tissues from the injection site or dilution with body fluids prior to gelation, leading to impaired formation and inhibited structural control after delivery¹⁴⁴.

In recent years, injectable cryogels have gained significant traction as preformed alternatives to in situ gelling materials (Fig 7). These cryogels are sponge-like, microporous assemblies that can be delivered via injection even after gelation is complete due to an inherent capability to collapse to a fraction of their hydrated size^38,145. Typically, these materials are formed by dissolving polymeric precursors and crosslinking initiators in water and freezing them at subzero temperatures. As these solutions freeze the dissolved solutes become concentrated in tightly confined non-frozen regions along grain boundaries where they can be incubated as polymerization takes place. Upon completion of the polymerization process, the material is heated to thaw the interspersed ice crystals and leaves behind fully formed hydrogels.

Figure 7. Cryogels form porous scaffolds with shape memory properties.

(A) Cryogels are formed when polymeric precursors are mixed with radical initiators and subsequently frozen in aqueous solution. This process confines the dissolved solutes to grain boundaries surrounding frozen water crystals where polymerization takes places. After thawing, the material forms a sponge-like hydrogel. (B) Cryogels exhibit shape memory properties. The hydrated materials can collapse to a fraction of their size as they are injected through standard needles, but subsequently rapidly revert to their pre-formed geometry after injection is complete. Cryogelation is easily adapted to form hydrogels of many different (C) sizes and (d) shapes. Source: Adapted with permission from Ref [38].

Cryogel materials have been successfully applied to a wide range of both cosmetic and regenerative therapies. In one study, Cheng et al. demonstrated that cryogels synthesized from methacrylated HA (cryo-MA-HA) could be utilized as improved fillers for soft-tissue reconstruction due to their relatively high stability in vivo¹⁴⁶. Pre-formed cryo-Me-HA gels injected subcutaneously in mice retained their shape for 30 days, remained situated at the injection site and demonstrated no adverse inflammatory effects over the course of the study. Additionally, skin-firmness, measured by durometer, showed no significant deviation between groups, implicating the materials ability to recapitulate desired mechanical properties at the area of injection. Regenerative metrics, including increased levels of cellularization and angiogenesis were detected throughout the filler relative to viscous HA and MA-HA controls indicating that the highly interconnected microporous structure of the material was sufficient to facilitate proper transport of nutrients and dissolved gasses and promote tissue incorporation. Lastly, the increased size of these preformed materials mitigated the risks associated with embolism stemming from accidental entry of accumulated debris into the surrounding vasculature. Together this data highlights several beneficial properties of cryogels, namely retaining injectability of a preformed microporous and deformable material while preserving enough mechanical integrity to mimic the mechanical properties of native tissues. Thus, the development of cryogels for dermal filling applications may result in highly stable off the shelf materials which can obviate the need for re-application in the form of sequential injections.

The highly interconnected micro-porosity of cryogels allows cells to easily migrate throughout the materials, making them an ideal environment for cellular delivery. To examine, the efficacy of such systems in vivo, Zeng et al. synthesized a gelatin microcryogel (GM) and seeded the material with hASCs¹⁴⁷. Evaluation of both differentiation markers and eluted small molecules after several days of culture revealed stark contrasts in cellular phenotype between cells seeded on GM and those grown on 2D substrate controls. hASCs grown within GM materials showed strongly upregulated expression of endothelial markers CD31, CD34 and HLA-DR while exhibiting inhibited expression for many osteogenic or adipogenic markers including RUNX2 and PPARγ even in the presence of induction medium. Additionally, secretion of regenerative growth factors such as VEGF, hepatocyte growth factor, BFGF, and PDGF showed higher expression at 48 hours in GM. The developed cryo-scaffolds also served as a protective barrier for seeded cells both by impeding absorption of cytotoxic solutes in the media and by buffering the shear forces of injection. During murine wound healing assays, GM hASC depots showed enhanced cell retention at the injection site as well as increased wound healing rates. In short, early application of cryogels towards wound healing are promising, but must be treated as preliminary until a wider base of literature provides support for these applications. Interestingly reports have also highlighted cryogels as a platform to modulate immune responses in vivo^148,149. These findings give hope that cryogels may one day be engineered to form micro-niches that both accurately coordinate stem cell response and promote regenerative outcomes by tuning overactive immune responses in chronic wounds.

4.3 Microporous Annealed Particle (MAP) Gels

Successful reformation of damaged tissues using hydrogel scaffolds is highly dependent on the ability of surrounding cells to penetrate and proliferate throughout the implanted biomaterial. Most scaffolds have accommodated this necessity through the incorporation of biodegradable crosslinks, which allow for cellular infiltration via material deterioration even if the scaffold is initially impermeable to cells. Although these approaches can often be implemented adequately, they are not robustly applicable. If degradation occurs too quickly, insufficient scaffolding will remain to support tissue ingrowth and macroscale cellular order will be lost. Conversely, degradation rates that are too slow will hinder tissue development and may result in fibrotic encapsulation of the implanted material. The heterogeneity of biological microenvironments between individuals coupled with the wide array of deteriorative stressors that hydrogels encounter in vivo underscore the difficulty in designing these materials for a “one size fits all” approach to treatment.

Our lab has developed a novel class of injectable biomaterial scaffolds, microporous annealed particle (MAP) gels that decouples the need for material degradation for tissue integration³⁷ (Fig 8). MAP gels are lattices of imperfectly annealed building blocks which form microporous networks that can be easily traversed by infiltrating cells. Generation of MAP building blocks occurs serially utilizing a microfluidic droplet generation device where two aqueous streams containing polymer backbone or cross-linker converge, mix thoroughly and are subsequently pinched into isolated droplets through the flow of a downstream oil phase. Synthesized droplets are subsequently collected, purified, and annealed into a microporous scaffold through secondary moieties present on the polymer backbone (Fig 8a). The reported system has the advantages of being both high throughput, with production rates of about 100 uL of pre-swollen microgel building blocks per hour in a single device, and readily customizable. To date MAP microgels have been synthesized from both PEG and HA backbones and have been annealed using a range of chemistries including, activated factor XIII mediated non-canonical amide linkages, light induced radical polymerization via Eosin Y, and through carbodiimide bonds from added PEG-NHS¹⁵⁰.

Figure 8. Microporous annealed particle hydrogels decouple tissue integration from material degradation.

(A) In the microporous annealed particle (MAP) gel system, microfluidically generated hydrogel beads are collected, washed and subsequently annealed via peptide functional groups present on their surface. This forms a microporous scaffold that allows cellular infiltration without degradation of the bonds that hold the material together. (B) Microgels are small enough to be injected through a standard syringe, and can conform to the shape of their injection cavity after annealing. (C) MAP hydrogels were evaluated for their effectiveness as regenerative scaffolds in a murine wound healing model, and demonstrated improved rates of wound closure versus no treatment, non-annealed microgels, and physically matched non-porous controls. Source: Adapted with permission from Ref [37].

Although MAP gels have demonstrated slight variations in both young's modulus and pore size dependent on the annealing chemistry chosen, these fluctuations do not hinder their utility as tissue culture scaffolds. In vitro, suspensions of human dermal fibroblasts, hASCs, and bone marrow-derived mesenchymal stem cells each exhibited high cell viability (greater than 93%) when cultured within the MAP scaffold for 24 hours. Furthermore, seeded cells began exhibiting a spread morphology within the first 90 minutes of culture and extended to form complex interconnected networks as early as 2 days after incorporation. This is in stark contrast to both physically and chemically matched non-porous controls which maintained cellular viability but did not allow the formation of an interconnected network.

In vivo studies complemented these findings by showing significantly increased rates of wound closure (39% after 7 days) for MAP scaffolds injected into BALB/c mice when compared to no treatment (19%) and non-porous control (7%) (Fig 8c). Subsequent immunohistochemical analysis of treated wounds revealed large scale stratified morphology among cells expressing epithelial markers keratin-5, keratin-14, and CD49f above the implanted gel. Notably, the scaffold was able to support the formation of complex structures such as hair follicles, supporting sebaceous glands and well-developed vasculature, indicated by the collective migration of endothelial cells and supporting pericytes throughout the material. Although this technology has demonstrated significant improvements for the treatment of dermal wounds, MAP gel is still in its infancy. In the future our lab hopes to develop building blocks with different sizes, elastic moduli, and functional groups to form scaffolds with three dimensional patterns or gradients of bioactive molecules. Studying the effect that these three-dimensional patterns have on tissue regeneration in vivo may enable the development of next generation scaffolds that can interface with the body for predictable and tunable regenerative responses as well as improved clinical outcomes.

5. Conclusion

Although numerous injectable hydrogels have been developed for drug delivery and wound repair, our ability to synthetically replicate the complexities of the native ECM are still unrefined at best. Ultimately, in order to develop broadly applicable hydrogels we must accomplish three goals. First, three-dimensional patterning of gradients within hydrogels, both in terms of bioactive functional groups and scaffold material properties, will have to be easily achievable if we have any hope of matching ECM variations between patients and among tissue types. Second, the presentation of integrins, growth factors, and small molecule drugs must occur in a dynamic but controllable manner such that coordinated cellular responses can be precisely induced with user control from the clinic. And lastly, materials must be synthesized which not only promote cellular adhesion, but which are also malleable enough to remain intact after considerable reorganization from infiltrating cellular populations. Here, we have highlighted many designs which begin to address these goals, however there is still much more work to be done.

Currently, clinical availability is limited for all but the most simple systems due to high production costs and lengthy regulatory approval processes¹⁵¹. However, the field of hydrogel design is rapidly evolving and may eventually translate to a library of single dose, minimally invasive, regenerative products. One particularly striking advance was reported by Brudno and coworkers, who demonstrated that an implanted drug delivery depot could regenerate its therapeutic payload after exhaustion simply through subsequent intravenous injection of the active molecule conjugated to oligonucleotide docking strands¹⁵². Such breakthroughs highlight how injectable hydrogels may eventually work synergistically with current clinical practice, to eliminate shortcomings, such as short half-lives or inefficient delivery to target sites in the case of multiple injections. As the burden of chronic wounds is felt most heavily among the elderly population, the demand for such injectable skin substitutes and drug delivery vehicles will only grow alongside the average life expectancy. Thus, injectable hydrogel biomaterials are primed to make a lasting impact on regenerative medicine for many years to come.